Radiation imaging apparatus and control method thereof, and radiation imaging system

ABSTRACT

An FPD detects an X-ray image of an object. The FPD includes a plurality of pixels arranged in its image capturing field. Each pixel receives X-rays emitted from an X-ray source, and outputs a pixel value in accordance with an X-ray dose applied thereto. A pixel determiner determines a minimum-value pixel out of the pixels based on the pixel values of the pixels. The minimum-value pixel is a pixel whose pixel value is the lowest. The pixel determiner sets the minimum-value pixel as an exposure control pixel. A comparator compares a first integrated value, which is an integrated value of the pixel values of the minimum-value pixel, with a predetermined first threshold value. The comparator performs X-ray emission control such that, when the first integrated value has reached the first threshold value, the X-ray source stops emitting the X-rays.

This application is a Continuation of U.S. application Ser. No.13/754,270, filed on Jan. 30, 2013, which claims priority under 35U.S.C. §119(a) to Application No. 2012-021868, filed in Japan on Feb. 3,2012, all of which are hereby expressly incorporated by reference intothe present application.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to a radiation imaging apparatus fortaking a radiographic image from radiation passed through an object, acontrol method of the radiation imaging apparatus, and a radiationimaging system having the radiation imaging apparatus.

2. Description Related to the Prior Art

In a medical field, a radiation imaging system, for example, an X-rayimaging system using X-rays is widely known. The X-ray imaging system isconstituted of an X-ray generating apparatus for applying the X-rays toan object (a body portion, for example, a chest of a patient), and anX-ray imaging apparatus for taking an X-ray image by reception of theX-rays passed through the object.

In recent years, the X-ray imaging apparatus that uses a flat paneldetector (FPD) as a detection panel, instead of an X-ray film or animaging plate (IP), becomes widespread. The FPD has a matrix of pixelseach of which produces and accumulates signal charge in accordance withan X-ray dose applied thereto. The FPD converts the signal charge of thepixels into a voltage signal by its signal processing circuit. Thereby,the FPD electrically detects the X-ray image, and outputs the X-rayimage as digital image data.

With the aim of reducing X-ray exposure of a patient and improving X-rayimage quality, some X-ray imaging systems have an automatic exposurecontrol (AEC) function for automatic control of an X-ray dose. Forexample, Japanese Patent Laid-Open Publication No. 09-055298 disclosesan X-ray fluoroscopic apparatus having a video camera, being an imagedetector, and an image intensifier disposed in front of the videocamera. In this apparatus, the image intensifier converts an X-ray imageinto an optical image, and the video camera captures a moving image tobe displayed on a monitor. The image intensifier is in the shape of arectangle, and is provided with three intensity detection sensors thatare disposed at the middle of an upper portion of the rectangle and atthe right and left of a lower portion of the rectangle, respectively, asX-ray sensors for detecting the X-ray dose. The video camera capturesthe moving image of an object in observation, so a fluoroscopic image isdisplayed on the monitor. During the display (fluoroscopy), pixel valuesof the fluoroscopic image are detected and a histogram of the pixelvalues is produced in each individual area corresponding to the positionof each X-ray sensor. Based on the histograms, an unexposed area towhich no X-ray is applied, an object area, and a directly exposed areato which the X-rays are directly applied without through the object aredetermined. Out of the three X-ray sensors, the one disposed in theobject area is chosen. In taking an X-ray image by use of an X-ray film,X-ray exposure time is controlled based on the X-ray dose detected bythe chosen X-ray sensor.

Also, U.S. Pat. No. 7,433,445 discloses a radiation imaging apparatusfor mammography. This apparatus includes an image detector composed ofan FPD, and an exposure control sensor disposed in a positioncorresponding to an outer edge of the image detector. In this apparatus,an X-ray dose (necessary dose) necessary for obtaining desirable imageequality is calculated from the thickness of an object, X-rayabsorptance, and the like, prior to performing mammography. During themammography, an X-ray dose (detected dose) detected by the exposurecontrol sensor is compared with the necessary dose. When the detecteddose has reached the necessary dose, X-ray emission is stopped.

The X-ray fluoroscopic apparatus of the Japanese Patent Laid-OpenPublication No. 09-055298 is provided with the plurality of X-raysensors for use in AEC. The X-ray sensors are used for identification ofthe object area, out of the three areas of the unexposed area, theobject area, and the directly exposed area. The X-ray sensor has lowspatial resolution, and its detection surface has fixed size and is in afixed position. Depending on the type, size, shape, or the like of abody portion, there may be cases where AEC cannot be performedappropriately, and desirable image quality cannot be obtained.

This is because, for example, in chest radiography, the object areaincludes lung fields, a mediastinum, and a diaphragm that have differentX-ray transmittances from each other. The difference in the X-raytransmittance causes variations in the X-ray dose to be transmitted.Therefore, the image quality differs depending on which part is used fordetecting the X-ray dose as a reference of AEC. In general, the higherthe density, the finer the graininess of an X-ray image would be. Thus,apart having a low X-ray transmittance is more preferably used fordetecting the X-ray dose than a part having a high X-ray transmittance,because increase in the density of the entire image facilitatesimproving the image quality.

However, in the Japanese Patent Laid-Open Publication No. 09-055298, theplurality of X-ray sensors are disposed in the fixed positions, and havethe fixed size and the low spatial resolution. Thus, the X-ray sensorsmay not be able to detect the X-ray dose at an appropriate position inthe object area, depending on the type, size, shape, or the like of thebody portion. Therefore, the apparatus of the Japanese Patent Laid-OpenPublication No. 09-055298 may fail to perform AEC appropriately,depending on the type, size, shape, or the like of the body portion.

On the other hand, in the U.S. Pat. No. 7,433,445, there is only oneexposure control sensor provided at the outer edge of the imagedetector, and its detection surface has fixed size and position. Thus,as with the Japanese Patent Laid-Open Publication No. 09-055298, theapparatus of the U.S. Pat. No. 7,433,445 may fail to perform appropriateAEC depending on the type, size, shape, or the like of the body portion,and fail to obtain favorable image quality.

Furthermore, both the X-ray sensors of the Japanese Patent Laid-OpenPublication No. 09-055298 and the exposure control sensor of the U.S.Pat. No. 7,433,445 are provided separately from the image detector(video camera or FPD), and hence may cause complex structure.

SUMMARY OF THE INVENTION

An object of the present invention is to provide a radiation imagingapparatus that has structure simpler than ever and can take aradiographic image of favorable image quality irrespective of a bodyportion to be imaged, a control method of the radiation imagingapparatus, and a radiation imaging system having the radiation imagingapparatus.

To achieve the above and other objects of the present invention, aradiation imaging apparatus according to the present invention includesan image detector, a pixel determiner, and a comparator. The imagedetector, which detects a radiographic image of an object, includes aplurality of pixels arranged in an image capturing field. Each of thepixels receives radiation emitted from a radiation source and outputs apixel value in accordance with a received radiation dose. The pixeldeterminer determines at least one typical low-value pixel from thepixels based on the pixel values, and sets the typical low-value pixelas an exposure control pixel. The comparator compares a first integratedvalue being an integrated value of the pixel value of the typicallow-value pixel with a predetermined first threshold value, and performsradiation emission control such that, when the first integrated valuehas reached the first threshold value, the radiation source stopsemitting the radiation.

The pixel determiner preferably determines at least one typicalhigh-value pixel from the pixels based on the pixel values, and sets thetypical high-value pixel as another exposure control pixel. Thecomparator preferably compares a second integrated value being anintegrated value of the pixel value of the typical high-value pixel witha predetermined second threshold value, and performs the radiationemission control such that, when the second integrated value has reachedthe second threshold value, the radiation source stops emitting theradiation even if the first integrated value has not reached the firstthreshold value.

The radiation imaging apparatus preferably further includes anirradiation field determiner for determining an irradiation field, whichis a field irradiated with the radiation in the image capturing field,based on the pixel values. The pixel determiner preferably determines inthe irradiation field a directly exposed area being an area applied withthe radiation directly without through the object, an implant area beingan area of an implant implanted in the object, and an object area beingan area excluding the directly exposed area and the implant area fromthe irradiation field. The typical low-value and high-value pixels arepreferably determined out of the pixels in the object area.

The pixel determiner preferably determines the object area based on ahistogram of the pixel values of the pixels in the irradiation field.

The pixel determiner may determine the typical low-value pixel out ofthe pixels present within an index area, which is predetermined in theimage capturing field in accordance with a body portion to be imaged.

The pixel determiner may determine the typical high-value pixel out ofthe pixels present within an interest area, which is predetermined inthe image capturing field in accordance with the body portion to beimaged.

It is preferable that radiation absorptance is higher in the index areathan in the interest area.

The pixels may include a plurality of normal pixels for specific use indetection of the radiographic image, and a plurality of detection pixelsdistributed throughout the image capturing field to detect the radiationdose.

The radiation imaging apparatus may further include a pixel valueestimator for estimating the pixel value of the normal pixel based onthe pixel values of the detection pixels near the normal pixel to beestimated. The pixel determiner preferably determines the typicallow-value and high-value pixels based on the estimated pixel values.

The image detector may have a plurality of pixel groups each includingone or more normal pixels and one or more detection pixels. Thedetection pixels are laid out differently between the pixel groupsadjoining to each other. The pixel value estimator may estimate thepixel value of the normal pixel of a first pixel group, based on thepixel value of the detection pixel belonging to the first pixel groupand the pixel value of the detection pixel belonging to a second pixelgroup adjoining to the first pixel group.

The pixel determiner may determine the typical low-value and high-valuepixels out of the detection pixels.

Signal lines electrically connected to the pixels may be routed in theimage capturing field to output the pixel values. The detection pixelmay be connected to the signal line directly or through a switchingelement.

The pixels may include a combined pixel that is composed of a firstsubpixel functioning as the normal pixel and a second subpixelfunctioning as the detection pixel.

A radiation imaging system of the present invention includes a radiationgenerating apparatus and a radiation imaging apparatus. The radiationgenerating apparatus includes a radiation source for emitting radiationto an object, and a source controller for controlling operation of theradiation source. The radiation imaging apparatus includes an imagedetector, a pixel determiner, and a comparator. The image detector,which detects a radiographic image of an object, includes a plurality ofpixels arranged in an image capturing field. Each of the pixels receivesthe radiation emitted from the radiation source, and outputs a pixelvalue in accordance with an applied radiation dose. The pixel determinerdetermines at least one typical low-value pixel from the pixels based onthe pixel values, and sets the typical low-value as an exposure controlpixel. The comparator compares a first integrated value being anintegrated value of the pixel value of the typical low-value pixel witha predetermined first threshold value, and performs radiation emissioncontrol such that, when the first integrated value has reached the firstthreshold value, the radiation source stops emitting the radiation.

A control method of a radiation imaging apparatus includes the steps ofdetermining at least one typical low-value pixel from pixels based onpixel values, and setting the typical low-value pixel as an exposurecontrol pixel; and comparing a first integrated value being anintegrated value of the pixel value of the typical low-value pixel witha predetermined first threshold value, and performing radiation emissioncontrol such that, when the first integrated value has reached the firstthreshold value, a radiation source stops emitting radiation.

According to the present invention, exposure control is carried outbased on the pixel value of the pixel in the image capturing field ofthe image detector. Thus, it is possible to simplify the structure ofthe radiation imaging apparatus. Furthermore, at least one typicallow-value pixel is determined out of the pixels of the image capturingfield, and the exposure control is performed based on the pixel value ofthe typical low-value pixel. Therefore, it is possible to obtain theradiographic image having favorable image quality, irrespective of thebody portion to be image.

BRIEF DESCRIPTION OF THE DRAWINGS

For more complete understanding of the present invention, and theadvantage thereof, reference is now made to the subsequent descriptionstaken in conjunction with the accompanying drawings, in which:

FIG. 1 is an explanatory view showing the schematic structure of anX-ray imaging system;

FIG. 2 is a block diagram of the X-ray imaging system;

FIG. 3 is a perspective view of an electronic cassette;

FIG. 4 is a block diagram showing the structure of an FPD;

FIG. 5 is a block diagram showing the structure of an exposurecontroller;

FIG. 6 is a flowchart of an irradiation field determination process;

FIG. 7 is a graph showing the correlation between a pixel value of anormal pixel and a pixel value of a short pixel;

FIG. 8 is an explanatory view showing an example of disposition of theshort pixels;

FIGS. 9A to 9D are explanatory views of the irradiation fielddetermination process;

FIGS. 10A and 10B are graphs showing an example of profiles that areproduced to obtain edges of the irradiation field;

FIG. 11 is a flowchart of a process of setting a minimum-value pixel anda maximum-value pixel;

FIG. 12 is a graph showing an example of a histogram that is produced todetermine an object area;

FIG. 13A is a graph showing an example of a profile in the absence of animplant;

FIG. 13B is a graph showing an example of a profile in the presence ofthe implant;

FIG. 14 is an explanatory view showing areas in which the minimum-valuepixel and the maximum-value pixel are located;

FIG. 15 is a flowchart of an AEC process using first and secondintegrated values;

FIG. 16 is a flowchart of an X-ray imaging process of an X-ray imagingsystem; and

FIG. 17 is an explanatory view showing the structure of a detectionpixel of another embodiment.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

As shown in FIG. 1, an X-ray imaging system 10 is constituted of anX-ray generating apparatus 11 for generating X-rays and an X-ray imagingapparatus 12 for taking an X-ray image from the X-rays passed through abody portion (object) of a patient H. The X-ray imaging apparatus 12includes an electronic cassette 13 for detecting the X-ray image and aconsole 14 that controls the electronic cassette 13 and performs imageprocessing of the X-ray image. In the X-ray imaging system 10, theconsole 14 is communicatably connected to the X-ray generating apparatus11 (concretely, a source controller 17) through a cable 15. Theelectronic cassette 13 and the console 14 are wirelessly communicatablewith each other. The X-ray imaging system 10 carries out AEC (automaticexposure control) in which the console 14 stops X-ray emission from theX-ray generating apparatus 11 at the instant when an X-ray dose detectedby the electronic cassette 13 has reached a predetermined value.

The X-ray generating apparatus 11 is constituted of an X-ray source 16,the source controller 17 for controlling the X-ray source 16, and anemission switch 18 for commanding the start of X-ray emission. The X-raysource 16 has an X-ray tube 16 a for emitting the X-rays and acollimator 16 b for limiting an irradiation field of the X-rays emittedfrom the X-ray tube 16 a. The X-ray tube 16 a has a cathode being afilament for emitting thermoelectrons and an anode (target) thatradiates the X-rays by collision of the thermoelectrons emitted from thecathode. The collimator 16 b is composed of, for example, four X-rayshielding lead plates disposed on each side of a rectangle so as to forman irradiation opening in its middle through which the X-rays propagate.A parallel shift of the lead plates varies the size of the irradiationopening to limit the irradiation field.

The electronic cassette 13 is detachably loaded in a holder of animaging stand 29 or an imaging table (not shown) in such a position thatan image capturing field 41 (see FIG. 4) of an FPD (image detector) 26is opposed to the X-ray source 16. The imaging stand 29 or the imagingtable may be designed specific to the electronic cassette 13, orsharable with a film cassette and an IP cassette. The electroniccassette 13 may be used by itself with being put on a bed under thepatient H lying, or held by the patient H himself/herself.

As shown in FIG. 2, the source controller 17 includes a high voltagegenerator 20, a control unit 21, and a wired communicator 22. The highvoltage generator 20 supplies high tube voltage to the X-ray source 16.The control unit 21 controls the tube voltage, tube current, irradiationtime, and the like. The wired communicator 22 establishes communicationwith the console 14. The tube voltage determines radiation quality(energy spectrum) of the X-rays emitted from the X-ray source 16. Thetube current determines the X-ray dose per unit of time. The highvoltage generator 20 converts input voltage into the high voltage by atransformer, and supplies drive power to the X-ray source 16 through ahigh voltage cable.

An imaging condition including the tube voltage, the tube current, theirradiation time, and the like is inputted from the console 14 to thecontrol unit 21 through the wired communicator 22. The control unit 21sets up a drive condition of the X-ray source 16 based on the imagingcondition. The imaging condition may be inputted from an operation panel23 provided in the source controller 17.

The emission switch 18 connected to the control unit 21 of the sourcecontroller 17 through a signal cable is operated by a radiologicaltechnician. The emission switch 18 is a two-step press switch, forexample. Upon a half press of the emission switch 18, a warm-up startsignal is issued to start warming up the X-ray source 16. Upon a fullpress, an emission start signal is issued to start the X-ray emissionfrom the X-ray source 16. The warm-up start signal and the emissionstart signal issued from the emission switch 18 are inputted to thecontrol unit 21 through the signal cable.

While the emission switch 18 is fully pressed, the X-ray source 16 isallowed to emit the X-rays. If the full press of the emission switch 18is released before performing AEC, the X-ray emission is stopped. Thus,it is possible to immediately stop the X-ray emission in case ofemergency.

Upon receiving the emission start signal from the emission switch 18,the control unit 21 of the source controller 17 issues an emission startcommand to the X-ray source 16, and makes the high voltage generator 20start electric power supply for the X-ray emission. Upon transmission ofthe emission stop signal from the electronic cassette 13 through theconsole 14 to the source controller 17 by AEC, the control unit 21issues an emission stop command to the X-ray source 16, and makes thehigh voltage generator 20 stop the electric power supply.

The electronic cassette 13 includes the FPD 26, an exposure controller32, a memory 33, a wireless communicator 34, a power source 35, and acontrol unit 36. The FPD 26 detects an X-ray image based on the X-raysapplied to its irradiation surface 25 through the object. The exposurecontroller 32 performs AEC. The memory 33 stores image data outputtedfrom the FPD 26. The wireless communicator 34 establishes communicationwith the console 14. The power source 35 supplies electric power from abattery to each part of the electronic cassette 13. The control unit 36controls the entire operation of the electronic cassette 13. Thecomponents described above are contained in a portable housing 27.

As shown in FIG. 3, the housing 27 is in the shape of a rectangular flatbox approximately the same size as the film cassette and the IPcassette. The housing 27 is provided with the battery for supplying theelectric power to the electronic cassette 13 on a surface opposite tothe irradiation surface 25.

The console 14 includes a monitor 75, an input device 76, and a mainbody 14 a. The monitor 75 displays an examination order, the X-rayimage, and the like. The input device 76 is used for input of theimaging condition and the like. The main body 14 a is constituted of animage storage 77 for storing the X-ray image data, a wired communicator78 for establishing communication with the source controller 17, awireless communicator 79 for establishing communication with theelectronic cassette 13, and a control unit 80 for controlling the entireoperation of the console 14.

The control unit 80 of the console 14 receives input of the examinationorder including information about the sex and age of the patient, a bodyportion to be imaged, an examination purpose, and the like, and displaysthe examination order on the monitor 75. The examination order isinputted from an external system that manages patient data orexamination data related to radiography such as a HIS (hospitalinformation system) or a RIS (radiography information system) connectedthrough the wired communicator 78 or the wireless communicator 79, orinputted manually by the radiological technician with the input device76. The radiological technician inputs the imaging condition, whichincludes the tube voltage, the tube current, the irradiation time, fromthe input device 76 to the control unit 80 based on the contents of theexamination order displayed on the monitor 75.

The control unit 80 transmits the imaging condition to the electroniccassette 13 and the source controller 17, and sets up in the electroniccassette 13 a signal processing condition of the FPD 26 including firstand second threshold values and the like. The control unit 80 receivesthe image data from the electronic cassette 13, and applies to the imagedata various types of image processing such as gamma correction andfrequency processing. The X-ray image after the image processing isdisplayed on the monitor 75 of the console 14. Also, the X-ray imagedata is written to the image storage 77 composed of a hard disk drive orthe like, and/or a data storage device such as an image storage serverconnected to the console 14 through a network.

As shown in FIG. 4, the FPD 26 has a TFT active matrix substrate, and isformed with the image capturing field 41 that is composed of a pluralityof pixels (normal pixels 40 and short pixels 55) arranged in the TFTactive matrix substrate. Each pixel produces signal charge in accordancewith an X-ray dose incident thereon. The plurality of pixels arearranged into a two-dimensional matrix with n rows (Y direction) and mcolumns (X direction) at a predetermined pitch. The FPD 26 also includesa gate driver 42 and a signal processing circuit 43. The gate driver 42drives the normal pixels 40 to control readout of the signal charge. Thesignal processing circuit 43 converts the signal charge read out fromthe normal pixels 40 into digital image data. The gate driver 42 and thesignal processing circuit 43 are controlled by the control unit 36.

The FPD 26 is of an indirect conversion type, which has a scintillator(not shown) for converting the X-rays into visible light. The pixelsperform photoelectric conversion of the visible light produced by thescintillator. The scintillator is disposed in front of the imagecapturing field 41 so as to be opposed to the entire image capturingfield 41 having an arrangement of the pixels. The scintillator is madeof phosphor such as CsI (cesium iodide) or GOS (gadolinium oxysulfide).Note that, a direct conversion type FPD, which has a conversion layer(amorphous selenium or the like) for directly converting the X-rays intoelectric charge, may be used instead.

The normal pixel 40 includes a photodiode 45, a capacitor (not shown),and a thin film transistor (TFT) 46. The photodiode 45 being aphotoelectric conversion element produces electric charge (electron andhole pairs) upon entry of the visible light. The capacitor accumulatesthe electric charge produced by the photodiode 45. The TFT 46 functionsas a switching element.

The photodiode 45 is composed of a semiconducting layer (of a PIN type,for example) of a-Si (amorphous silicon) or the like, and upper andlower electrodes disposed on the top and bottom of the semiconductinglayer. The lower electrode of the photodiode 45 is connected to the TFT46. The upper electrode of the photodiode 45 is connected to a bias line(not shown).

Through the bias line, bias voltage is applied to the upper electrode ofthe photodiode 45 of every pixel in the image capturing field 41. Sincethe application of the bias voltage produces an electric field in thesemiconducting layer of the photodiode 45, the electric charge (electronand hole pairs) produced in the semiconducting layer by thephotoelectric conversion is attracted to the upper and lower electrodes,one of which has positive polarity and the other has negative polarity.Thereby, the electric charge is accumulated in the capacitor.

A gate electrode of the TFT 46 is connected to a scan line 48. A sourceelectrode of the TFT 46 is connected to a signal line 49. A drainelectrode of the TFT 46 is connected to the photodiode 45. The scanlines 48 and the signal lines 49 are routed into a lattice. The numberof the scan lines 48 coincides with the number of the rows (n rows) ofthe pixels provided in the image capturing field 41, and all the pixelsarranged in the same row are connected to the same scan line 48. Thenumber of the signal lines 49 coincides with the number of the columns(m columns) of the pixels, and all the pixels arranged in the samecolumn are connected to the same signal line 49. Every scan line 48 isconnected to the gate driver 42, and every signal line 49 is connectedto the signal processing circuit 43.

The gate driver 42 drives the TFTs 46 so that the FPD 26 carries out acharge accumulation operation, a readout operation, and a resetoperation. In the charge accumulation operation, the normal pixels 40accumulate the signal charge by an amount corresponding to the X-raydose incident thereon during the X-ray emission. In the readoutoperation, the signal charge is read out from the normal pixels 40 afterthe X-ray emission. The reset operation is performed immediately beforethe X-ray emission to discharge and reset the signal charge accumulatedin the normal pixels 40. The control unit 36 controls start timing ofeach operation described above carried out by the gate driver 42.

In the charge accumulation operation, every TFT 46 is turned off, soevery normal pixel 40 accumulates the signal charge. In the readoutoperation, the gate driver 42 sequentially issues gate pulses G1 to Gneach of which drives the TFTs 46 of the same row at a time. Thereby, thescan lines 48 are activated one by one so as to turn on the TFTs 46connected to the activated scan line 48 on a row-by-row basis.

Upon turning on the TFTs 46 of the single row, the signal chargeaccumulated in the normal pixels 40 of the single row is inputted to thesignal processing circuit 43 through the signal lines 49. The signalprocessing circuit 43 reads out output voltage corresponding to thesignal charge as voltage signals D1 to Dm. Each of the analog voltagesignals D1 to Dm is converted into a digital pixel value, being adetection value of each pixel, after predetermined gain adjustment. Thispixel value is also called QL (quantum level) value. The image data thatis composed of the pixel values of the pixels is outputted to the memory33 contained in the electronic cassette 13.

Dark current occurs in the semiconducting layer of the photodiode 45irrespective of the presence or absence of entry of the X-rays. Darkcharge of the dark current is accumulated in the capacitor due to theapplication of the bias voltage. The dark charge becomes noise of theimage data, and therefore the reset operation is carried out to removethe dark charge. The reset operation is an operation of dischargingunnecessary electric charge e.g. the dark charge accumulated in thenormal pixels 40 through the signal lines 49.

The reset operation adopts a sequential reset method, for example, bywhich the pixels are reset on a row-by-row basis. In the sequentialreset method, as in the case of the readout operation of the signalcharge, the gate driver 42 sequentially issues the gate pulses G1 to Gnto the scan lines 48 to turn on the TFTs 46 of the pixels on arow-by-row basis. While the TFT 46 is turned on, the dark charge flowsfrom the pixel through the signal line 49 into the signal processingcircuit 43.

In the reset operation, in contrast to the readout operation, outputvoltage corresponding to the dark charge is not read out. Insynchronization with the issue of each of the gate pulses G1 to Gn, thecontrol unit 36 outputs a reset pulse RST to the signal processingcircuit 43. In the signal processing circuit 43, the input of the resetpulse RST turns on reset switches 51 a of integration amplifiers 51described later on, so the integration amplifiers 51 are reset.

Instead of the sequential reset method, a parallel reset method or anall pixels reset method may be used. In the parallel reset method, aplurality of rows of the pixels are grouped together, and sequentialreset is carried out in each group, so as to concurrently discharge thedark charge from the rows of the number of the groups. In the all pixelsreset method, the gate pulse is inputted to every row to concurrentlydischarge the dark charge from every pixel. Adoption of the parallelreset method and the all pixels reset method can reduce time requiredfor the reset operation.

The signal processing circuit 43 is provided with the integrationamplifiers 51, a MUX 52, an A/D converter 53, and the like. Oneintegration amplifier 51 is connected to each signal line 49. Theintegration amplifier 51 includes an operational amplifier and acapacitor connected between input and output terminals of theoperational amplifier. The signal line 49 is connected to one of twoinput terminals of the operational amplifier. The other input terminalof the operational amplifier is connected to a ground (GND). Theintegration amplifiers 51 integrate the signal charge inputted from thesignal lines 49, and convert the signal charge into the voltage signalsD1 to Dm, and output the voltage signals D1 to Dm.

In the readout operation for reading out the signal charge from everynormal pixel 40 after the charge accumulation operation, the TFTs 46 areturned on from row to row by the gate pulses. The signal charge flowsfrom the capacitors of the normal pixels 40 in the activated row intothe integration amplifiers 51 through the signal lines 49.

An output terminal of the integration amplifier 51 of every column isconnected to the MUX 52 through another amplifier (not shown) foramplifying each of the voltage signals D1 to Dm and a sample holder (notshown) for holding each of the voltage signals D1 to Dm. The MUX 52sequentially chooses one of the plurality of integration amplifiers 51connected in parallel, in order to input the voltage signals D1 to Dm inseries from the chosen integration amplifier 51 to the A/D converter 53.

The A/D converter 53 converts the analog voltage signals D1 to Dm intothe digital pixel values in accordance with their signal levels, andoutputs the pixel values to the memory 33. The pixel values are storedin the memory 33 as the X-ray image data representing the X-ray imagewith being associated with the coordinates of each normal pixel 40 inthe image capturing field 41.

After the voltage signals D1 to Dm of one row are outputted from theintegration amplifiers 51, the control unit 36 issues the reset pulseRST to the integration amplifiers 51 to turn on the reset switches 51 aof the integration amplifiers 51. Thus, the signal charge of the singlerow that is accumulated in the integration amplifiers 51 is reset. Afterthe reset of the integration amplifiers 51, the gate driver 42 outputsthe gate pulse of the next row, so the signal charge is read out fromthe normal pixels 40 of the next row. The above operation is repeated insequence to read out the signal charge from the normal pixels 40 ofevery row.

After completion of the readout from every row, the image datarepresenting the X-ray image of one frame is written to the memory 33.The control unit 36 applies image correction processing including offsetcorrection, sensitivity correction, and defect correction to this imagedata. In the offset correction, an offset component being fixed patternnoise caused by the individual difference and the environment of the FPD26 is eliminated. In the sensitivity correction, variations insensitivity among the photodiodes 45, variations in output property ofthe signal processing circuit 43, and the like are corrected. In thedefect correction, the pixel value of a defect pixel is linearlyinterpolated using the pixel value of the normal pixel 40 adjoining tothe defect pixel, based on defect pixel information produced in shipmentor periodic inspection. The pixel value of the short pixel 55 used inAEC, as described later on, is also subjected to the defect correction.The image data is read out from the memory 33, and is transmitted to theconsole 14 through the wireless communicator 34.

The FPD 26 is provided with not only the normal pixels 40 for specificuse in detection of the X-ray image but also the plurality of shortpixels 55, which are hatched in FIG. 4, in its image capturing field 41.The short pixel 55 is a detection pixel that detects an X-ray doseapplied to the FPD 26 through the object. The short pixels 55 are usedin AEC performed by the exposure controller 32 and obtainment ofoperation switching timing of the FPD 26.

The short pixels 55 are distributed evenly across the entire imagecapturing field 41 without being localized. The short pixels 55 occupy,for example, about 0.01% of the all pixels including the normal pixels40 and the short pixels 55 in the image capturing field 41. Thepositions of the short pixels 55 are known in manufacturing the FPD 26.The FPD 26 has a nonvolatile memory (not shown) that stores the position(coordinates) of every short pixel 55. The layout, number, and rate ofthe short pixels 55 are appropriately changeable.

Each short pixel 55 has the photodiode 45 and the TFT 46, as with thenormal pixel 40. The photodiode 45 of the short pixel 55 produces thesignal charge in accordance with the X-ray dose incident thereon. Thedifference in the structure between the short pixel 55 and the normalpixel 40 is that the short pixel 55 has a connection 55 a that brings ashort between the source and the drain of the TFT 46, and hence theshort pixel 55 has no switching function of the TFT 46. Thus, the signalcharge produced in the photodiode 45 of the short pixel 55 continuouslyflows out through the signal line 49 into the integration amplifier 51.Note that, instead of connecting the source and the drain of the TFT 46of the short pixel 55, the short pixel 55 may not be provided with theTFT 46 and the photodiode 45 may be directly connected to the signalline 49.

The signal charge from the short pixels 55 that has been inputted to theintegration amplifiers 51 is outputted to the A/D converter 53, as withthe signal charge from the normal pixels 40. The A/D converter 53converts the signal charge into digital pixel values Vout, and outputsthe pixel values Vout to the memory 33. The pixel values Vout of theshort pixels 55 are stored in the memory 33 with being associated withthe coordinates of each short pixel 55. Thus, the X-ray dose that hasbeen applied to each short pixel 55 is detected. The FPD 26 repeats thissampling operation of the pixel values Vout of the short pixels 55 at apredetermined rate during the X-ray emission. The exposure controller 32reads out the sampled pixel values Vout of the short pixels 55 from thememory 33 to carry out AEC.

As shown in FIG. 5, the exposure controller 32 is provided with anirradiation field determiner 57, a pixel determiner 58, a comparator 59,and an emission start/stop detector 60. The irradiation field determiner57, which determines the irradiation field of the X-rays applied fromthe X-ray source 16 in AEC, is composed of the short pixels 55 describedabove and a pixel value estimator 61. An irradiation field determinationprocess by the irradiation field determiner 57 will be hereinafterdescribed with referring to FIGS. 6 to 10B.

The pixel value estimator 61 estimates the pixel value of the normalpixel 40 based on the pixel value Vout of the short pixel 55 positionedin the vicinity of the normal pixel 40 (S10). As shown in FIG. 7, due tothe positive correlation between the pixel value of the normal pixel 40and the pixel value Vout of the short pixel 55 positioned in thevicinity thereof, the pixel value of the normal pixel 40 can beestimated from the pixel value Vout of the short pixel 55 with highaccuracy. Note that, the pixel value of the normal pixel 40 and thepixel value Vout of the short pixel 55 have a linear correlation in FIG.7, but may have a nonlinear correlation depending on the structure andpositions of the pixels.

The pixel value estimator 61 may estimate the pixel value of the normalpixel 40 based on the pixel value Vout of the short pixel 55 obtained bysingle sampling, or based on an integrated value of the pixel valuesVout, which are sampled two or more times from the single short pixel 55and integrated from one coordinate to another.

In this embodiment, to improve accuracy in the estimation of the pixelvalue of the normal pixel 40, a plurality of pixel groups each of whichhas a fixed number of pixels in row and column directions areestablished in the FPD 26. Each pixel group includes a predeterminednumber of normal pixels 40 and a predetermined number of short pixels55. The positions of the short pixels 55 differ between the pixel groupsadjoining in the column direction, for example. In an example shown inFIG. 8, first to sixth pixel groups 63 to 68, each including 3 by 3pixels in the row and column directions (X and Y directions), areestablished. In the first pixel group 63, the three short pixels 55 arearranged in a left column. In the second pixel group 64 adjoining to thefirst pixel group 63 in the column direction, the three short pixels 55are arranged in a middle column. In the third pixel group 65 adjoiningto the second pixel group 64 in the column direction, the three shortpixels 55 are arranged in a right column.

Taking a case of estimating the pixel values of the normal pixels 40arranged in a middle column of the first pixel group 63 as an example,the pixel value estimator 61 estimates the pixel values based on thepixel values of the short pixels 55 belonging to the same first pixelgroup 63 and the pixel values of the short pixels 55 belonging to thesecond pixel group 64. The pixel values of the normal pixels 40 arrangedin a right column of the first pixel group 63 are estimated based on thepixel values of the short pixels 55 belonging to the same first pixelgroup 63 and the pixel values of the short pixels 55 belonging to thethird pixel group 65. In a like manner, the pixel value of each of thenormal pixels 40 of the second to sixth pixel groups 64 to 68 isestimated from the pixel values of the short pixels 55 belonging to thesame pixel group and the pixel values of the short pixels 55 belongingto the near pixel group.

Note that, the pixel value of the normal pixel 40 may be estimated onlyfrom the pixel values of the short pixels 55 belonging to the same pixelgroup, without using the pixel values of the short pixels 55 belongingto the different pixel group. In this case, estimation processingbecomes easier, though its accuracy possibly decreases.

The irradiation field determiner 57 differentiates an image that iscomposed of the estimated pixel values of the normal pixels 40 toproduce a differential image, and obtains a barycenter of differentialvalues from the differential image (S11). In an example shown in FIGS.9A to 9D, the irradiation field is determined from an image of a righthand of the patient. As shown in FIG. 9A, the irradiation fielddeterminer 57 produces a differential image 70 by differentiation of theimage composed of the estimated pixel values of the normal pixels 40,and obtains a barycenter G from the differential image 70.

Next, potential points of the edge of the irradiation field areextracted (S12). As shown in FIG. 9B, the irradiation field determiner57 performs differential processing of an image 72, which is composed ofthe estimated pixel values of the normal pixels 40, with respect to aplurality of directions radiating from the barycenter G. To be morespecific, as shown in FIG. 10A, a profile of the pixel values of theimage 72 is produced in each radiating direction, and this profile issubjected to the differential processing to create a profile of absolutevalues of the differential values, as shown in FIG. 10B. Then, theabsolute differential values are compared with a predetermined thresholdvalue TH. The coordinates of the pixel that has the absolutedifferential value larger than the threshold value TH are extracted asthe potential point of the edge of the irradiation field.

Six radiating directions thirty degrees apart from each other are set inFIG. 9B, but the number of the radiating directions is preferablyincreased. In the above embodiment, the profile is produced in theplurality of directions radiating from the barycenter G of thedifferential image 72. However, if it is conceivable that therectangular irradiation field is not rotated or skewed with respect tothe image, an easier method is adoptable in which the profile may bedetected in two directions, that is, vertical and horizontal directions.Note that, other irradiation field determination methods are known inJapanese Patent No. 2525652, Japanese Patent Laid-Open Publication Nos.63-259538 and 10-162156, and the like. The methods described in the artmay be adopted instead.

Next, several edge points are determined from the extracted potentialpoints (S13). To be more specific, for example, the irradiation fielddeterminer 57 re-evaluates whether or not each midpoint between thepotential points next to each other is actually in the edge of theirradiation field, and determines eight edge points E1 to E8, as shownin FIG. 9C. In the next step, the irradiation field is determined fromthe edge points E1 to E8 (S14). As shown in FIG. 9C, the irradiationfield determiner 57 creates a polygonal field F1 by connecting thedetermined eight edge points E1 to E8 by straight lines. Then, as shownin FIG. 9D, the irradiation field determiner 57 corrects the position ofthe edge point E3 that is unnaturally recessed, and corrects the shapeof the field F1 in accordance with the shape of the irradiation fielddefined by the collimator 16 b to determine a rectangular irradiationfield F2.

The pixel determiner 58 determines pixels to be used in AEC, out of thenormal pixels 40 located within the irradiation field F2 determined bythe irradiation field determiner 57. Based on the pixel values of thenormal pixels 40 estimated by the pixel value estimator 61, the pixeldeterminer 58 determines a minimum-value pixel whose pixel value is thelowest and a maximum-value pixel whose pixel value is the highest. Theminimum-value pixel and the maximum-value pixel are used in AEC. Theminimum-value pixel is set as a typical low-value pixel that is used forobtaining the X-ray image with favorable image quality. Themaximum-value pixel is set as a typical high-value pixel that is usedfor preventing excessive X-ray exposure of the patient.

Referring to FIG. 11, a setting process of the minimum-value pixel andthe maximum-value pixel by the pixel determiner 58 will be described. Asshown in FIG. 12, the pixel determiner 58 creates a histogram of theestimated pixel values of the normal pixels 40 located within theirradiation field F2 (S20). The pixel determiner 58 analyzes the createdhistogram to obtain the minimum-value pixel and the maximum-value pixel.In the histogram, a horizontal axis represents the amount of theestimated pixel value, and a vertical axis represents the frequency ofappearance of each estimated pixel value. Then, the pixel determiner 58determines from the histogram, a directly exposed area to which theX-rays are directly applied without through the object and an implantarea in which an implant is present in the object. The pixel determiner58 determines an object area by excluding the directly exposed area andthe implant area from the irradiation field F2 (S21).

The pixel value is higher in the directly exposed area than in theobject area, because the X-rays are not absorbed by the object in thedirectly exposed area. The pixel determiner 58 detects a maximum peak,which represents the maximum estimated pixel value, out of peaks of thehistogram, for example. The pixel value of the maximum peak ismultiplied by a certain rate less than one, and the multiplied pixelvalue is set as a directly exposed area threshold value. In theirradiation field F2, an area having the pixel values that are equal toor more than the directly exposed area threshold value is determined asthe directly exposed area. On the other hand, when the object is aliving human body, the X-ray absorptance of the implant is higher thanthat of the object. Thus, the pixel value is lower in the implant areathan in the object area. The pixel determiner 58 detects a minimum peak,which represents the minimum estimated pixel value, out of the peaks ofthe histogram, for example. The pixel value of the minimum peak ismultiplied by a certain rate more than one, and the multiplied pixelvalue is set as an implant area threshold value. In the irradiationfield F2, an area having the pixel values that are equal to or less thanthe implant area threshold value is determined as the implant area. Notethan, the rates used for calculation of the directly exposed areathreshold value and the implant area threshold value are changeable inaccordance with the body portion to be imaged.

The pixel determiner 58 determines based on the histogram a pixel whoseestimated pixel value is the lowest (MIN) out of the normal pixels 40 inthe object area, and sets this pixel as the minimum-value pixel. In alike manner, the pixel determiner 58 determines a pixel whose estimatedpixel value is the highest (MAX), and sets this pixel as themaximum-value pixel (S22).

The directly exposed area and the implant area are excluded from theirradiation field F2, for the purpose of setting the minimum-value pixeland the maximum-value pixel to be used in AEC within the object areaexcluding the directly exposed area and the implant area. FIGS. 13A and13B show profiles of an irradiation field in a lateral direction inchest radiography. FIG. 13A is in the case of the absence of theimplant, while FIG. 13B is in the case of the presence of the implant inthe middle of the object. In FIGS. 13A and 13B, a horizontal axisrepresents the horizontal position of the normal pixels 40 in theirradiation field F2, and a vertical axis represents a pixel value. Asis known from the profiles, the pixel value is lower in the implant areathan in the object area. The pixel value is higher in the directlyexposed area than in the object area. Therefore, if the minimum andmaximum values are chosen from the pixel values of all the normal pixels40 located within the irradiation field F2, a pixel within the implantarea is set as the minimum-value pixel, and a pixel within the directlyexposed area is set as the maximum-value pixel. For this reason, in thisembodiment, the object area is determined by excluding the directlyexposed area and the implant area from the irradiation field F2, and theminimum-value pixel and the maximum-value pixel are set from the normalpixels 40 within the object area.

In the case of setting the minimum-value pixel and the maximum-valuepixel from the one-dimensional profile shown in FIG. 13B, the positionsof the minimum value (MIN) and the maximum value (MAX) correspond to thepositions of the minimum-value pixel and the maximum-value pixel. Sincethe object area is a two-dimensional area, there may be cases where theprofile of FIG. 13B includes either or neither of the MIN and MAX, as amatter of course.

In the case of the chest radiography, for example, the object areaincludes right and left lung fields, a mediastinum, a diaphragm, and thelike. In the object area, the lung fields have the highest X-raytransmittance, while the mediastinum and the diaphragm have the lowestX-ray transmittance. Therefore, in the chest radiography, as shown inFIG. 14, the minimum-value pixel having the MIN is set in an area 73 bcorresponding to the mediastinum and the diaphragm, while themaximum-value pixel having the MAX is set in areas 73 a corresponding tothe right and left lung fields. The coordinate data of the minimum-valuepixel and the maximum-value pixel is inputted to the comparator 59 andthe pixel value estimator 61.

The comparator 59 performs AEC based on the estimated pixel values ofthe minimum-value and maximum-value pixels set by the pixel determiner58. After the pixel determiner 58 sets the minimum-value andmaximum-value pixels, the pixel value estimator 61 estimates the pixelvalues of the minimum-value and maximum-value pixels, and the estimatedpixel values are inputted to the comparator 59 (see FIG. 5). Asdescribed above, the pixel values Vout of the short pixels 55 aresampled repeatedly at a predetermined rate. Whenever the sampling iscarried out, the pixel value estimator 61 estimates the pixel value ofeach of the minimum-value and maximum-value pixels based on the pixelvalues Vout of the short pixels 55, and inputs the estimated pixelvalues to the comparator 59. The estimated pixel value that iscalculated based on the pixel values Vout obtained by the singlesampling corresponds to an X-ray dose applied to the minimum-value ormaximum-value pixel per unit of time. The comparator 59 integrates theestimated pixel values of the minimum-value pixel inputted in eachsampling cycle, to calculate a first integrated value being anintegrated value of the estimated pixel values of the minimum-valuepixel. The comparator 59 also integrates the estimated pixel values ofthe maximum-value pixel inputted in each sampling cycle, to calculate asecond integrated value being an integrated value of the estimated pixelvalues of the maximum-value pixel. The first integrated valuecorresponds to a cumulative amount of the X-ray dose applied to theminimum-value pixel. The second integrated value corresponds to acumulative amount of the X-ray dose applied to the maximum-value pixel.

As shown in FIG. 15, a first threshold value to be compared with thefirst integrated value and a second threshold value to be compared withthe second integrated value are set in the comparator 59 (S30). Thecomparator 59 stores a plurality of types of first and second thresholdvalues, which correspond to the body portion to be imaged, in its memory(not shown), and chooses the first and second threshold values based onthe body portion included in the imaging condition transmitted from theconsole 14. The first threshold value represents a necessary doserequired for obtaining the favorable image quality of the X-ray image.The second threshold value represents a regulation value to preventexcessive X-ray exposure of the object.

The comparator 59 compares the second integrated value with the secondthreshold value (S31) to check whether or not the X-ray dose applied tothe object has reached the regulation value. In a case where the secondintegrated value is the second threshold value or more (YES in S31), thecomparator 59 judges that the applied X-ray dose has reached theregulation value, and issues an emission stop signal to the control unit36 of the electronic cassette 13 (S32). The emission stop signal istransmitted from the control unit 36 to the source controller 17 throughthe console 14. Thereby, the X-ray emission from the X-ray source 16 isstopped.

In a case where the second integrated value is less than the secondthreshold value (NO in S31), the comparator 59 compares the firstintegrated value with the first threshold value (S33) to check whetheror not the applied X-ray dose has reached the necessary dose. Thecomparator 59 repeats the comparison between the second integrated valueand the second threshold value and between the first integrated valueand the first threshold value (S31 and S33), until the first integratedvalue comes to the first threshold value or more. When the firstintegrated value is the first threshold value or more (YES in S33), thecomparator 59 judges that the applied X-ray dose has reached thenecessary dose, and issues the emission stop signal to the control unit36 of the electronic cassette 13 (S32). As described above, when thesecond integrated value has reached the second threshold value and theapplied X-ray dose has come to the regulation value (YES in S31), theX-ray emission from the X-ray source 16 is stopped, even if the firstintegrated value is less than the first threshold value.

As described above, AEC is carried out based on the first integratedvalue of the minimum-value pixel in the object area, so it is possibleto obtain the X-ray image with the favorable image quality. The higherthe density, the finer the graininess and the higher the image qualityof the X-ray image would be. In this embodiment, the minimum-value pixelthat is located in a part of the object area having the lowest X-raytransmittance is used as reference of AEC. Thus, the necessary dose iscertainly applied not only to the minimum-value pixel but also to theentire object area, so it is possible to obtain the X-ray image havingthe favorable image quality in the entire object area.

Also, since AEC is carried out based on the second integrated value ofthe maximum-value pixel in the object area, it is possible to preventexcessive X-ray exposure throughout the object area. For example, if apixel of the object area having relatively low X-ray transmittance isused as reference, and an integrated value of this reference pixel iscompared with the second threshold value (regulation value) forprevention of the excessive X-ray exposure, the X-ray dose could exceedthe regulation value in a part having the X-ray transmittance higherthan that of the reference pixel. In this embodiment, the maximum-valuepixel that is located in a part of the object area having the highestX-ray transmittance is used as reference of AEC. Thus, the applied X-raydose is the second integrated value or less in the entire object area,so it is possible to prevent the excessive X-ray exposure in the entireobject area.

The emission start/stop detector 60 monitors the pixel value Vout of theshort pixel 55 during the reset operation of the FPD 26 before the startof X-ray emission. The pixel value Vout of the short pixel 55 is sampledrepeatedly at a predetermined sampling rate during the reset operation.Whenever the sampling is performed, the pixel value Vout is inputted tothe emission start/stop detector 60. The emission start/stop detector 60compares the pixel value Vout with a predetermined emission startthreshold value. When the pixel value Vout has reached the emissionstart threshold value, the emission start/stop detector 60 judges thatthe X-ray source 16 has started the X-ray emission, and issues anemission start detection signal to the control unit 36. Upon receivingthe emission start detection signal, the control unit 36 shifts theoperation of the FPD 26 from the reset operation to the chargeaccumulation operation.

The sampling of the pixel value Vout of the short pixel 55 is continuedduring the X-ray emission. During the X-ray emission, the pixel valueVout is inputted to the emission start/stop detector 60, in addition tobeing used for AEC, as described above. The emission start/stop detector60 compares the pixel value Vout of the short pixel 55 with apredetermined emission stop threshold value during the X-ray emission.When the pixel value Vout has reached the emission stop threshold value,the emission start/stop detector 60 judges that the X-ray source 16 hasstopped the X-ray emission, and issues an emission stop detection signalto the control unit 36. Upon receiving the emission stop detectionsignal, the control unit 36 shifts the operation of the FPD 26 from thecharge accumulation operation to the readout operation.

Referring to FIG. 16, the operation of the X-ray imaging system 10 willbe described. The position adjustment is carried out among the imagingstand 29 loaded with the electronic cassette 13, the body portion of thepatient H, and the irradiation position of the X-ray source 16. Theexamination order including the sex and age of the patient, the bodyportion to be imaged, the examination purpose, and the like is inputtedto the console 14, and the imaging condition is set up based on theexamination order (S101). The console 14 transmits the imaging conditionto the electronic cassette 13 and the source controller 17.

The control unit 21 of the source controller 17 sets up a drivingcondition of the X-ray source 16 based on the imaging condition receivedfrom the console 14 (S301). The control unit 36 of the electroniccassette 13 sets up the first and second threshold values describedabove based on the imaging condition received from the console 14(S201).

The console 14 transmits a preparation command signal, which commandspreparation for imaging, to the electronic cassette 13 (S102). Uponreceiving the preparation command signal, the electronic cassette 13shifts the FPD 26 to a standby state (S202). In the standby state, theFPD 26 starts the reset operation, and the emission start/stop detector60 of the exposure controller 32 begins detecting the start of X-rayemission.

When the emission start signal is inputted from the emission switch 18,the source controller 17 issues the emission start command to the X-raysource 16 (S302). The X-ray source 16 starts applying the X-rays to theobject. The emission start/stop detector 60 compares the pixel valueVout of the short pixel 55 with the emission start threshold value. Whenthe pixel value Vout has reached the emission start threshold value, thestart of X-ray emission is detected (S203). Upon detecting the start ofX-ray emission, the TFTs 46 of the normal pixels 40 are turned off tostart the charge accumulation operation of the FPD 26 (S204).

The exposure controller 32 carries out AEC based on the pixel valuesVout of the short pixels 55 (S205). In AEC, the irradiation fielddeterminer 57 determines the irradiation field based on the pixel valuesVout of the short pixels 55. To be more specific, as shown in FIG. 6,the pixel value estimator 61 estimates the pixel value of every normalpixel 40 arranged in the image capturing field 41 based on the pixelvalues Vout of the short pixels 55. The irradiation field determiner 57determines the irradiation field based on the estimated pixel values ofthe normal pixels 40. Then, as shown in FIG. 11, the pixel determiner 58determines the directly exposed area, the implant area, and the objectarea based on the estimated pixel values of the normal pixels 40 locatedwithin the irradiation field. The pixel determiner 58 chooses the normalpixels 40 having the minimum estimated pixel value and the maximumestimated pixel value from the normal pixels 40 located within theobject area. The chosen normal pixel 40 having the minimum estimatedpixel value is set as the minimum-value pixel. The chosen normal pixel40 having the maximum estimated pixel value is set as the maximum-valuepixel.

The coordinate data of the minimum-value pixel and the maximum-valuepixel is inputted to the comparator 59 and the pixel value estimator 61.Whenever the pixel values Vout of the short pixels 55 are sampled, thepixel value estimator 61 estimates the pixel values of the minimum-valueand maximum-value pixels based on the pixel values Vout, and inputs theestimated pixel values to the comparator 59. The comparator 59integrates the estimated pixel values to obtain the first integratedvalue of the minimum-value pixel and the second integrated value of themaximum-value pixel. Also, as shown in FIG. 15, the first and secondthreshold values are set up in the comparator 59.

Referring to FIG. 15, the comparator 59 compares the second integratedvalue with the second threshold value, and compares the first integratedvalue with the first threshold value. The comparator 59 repeats thecomparison between the first integrated value and the first thresholdvalue, until the second integrated value reaches the second thresholdvalue. When the first integrated value has reached the first thresholdvalue, the comparator 59 issues the emission stop signal. When thesecond integrated value has reached the second threshold value, thecomparator 59 issues the emission stop signal, even if the firstintegrated value has not reached the first threshold value.

The emission stop signal is transmitted from the electronic cassette 13to the source controller 17 through the console 14 (S103). Uponreceiving the emission stop signal, the source controller 17 issues theemission stop command to the X-ray source 16 to stop the X-ray emission(S303).

The emission start/stop detector 60 of the exposure controller 32compares the pixel value Vout of the short pixel 55 with the emissionstop threshold value. When the pixel value Vout has reached the emissionstop threshold value, the stop of X-ray emission is detected (S206).Upon detecting the stop of X-ray emission, the FPD 26 stops the chargeaccumulation operation and starts the readout operation (S207). TheX-ray image data read out from the FPD 26 is transmitted from theelectronic cassette 13 to the console 14 (S208). The X-ray image data issubjected to the predetermined image processing, and written to theimage storage 77 (S104).

The emission start/stop detector 60 is provided in this embodiment sothat the FPD 26 detects the start and stop of X-ray emission, but theemission start/stop detector 60 may be omitted. In the case of theabsence of the emission start/stop detector 60, the X-ray generatingapparatus 11 transmits the start and stop of X-ray emission to the FPD26 by electrical communication.

As described above, the minimum-value pixel having the lowest pixelvalue is chosen from the pixels in the image capturing field 41, and thefirst integrated value of the minimum-value pixel is compared with thenecessary dose in AEC. Since the pixel is used as an AEC sensor formeasuring the X-ray dose, it is possible to obtain higher spatialresolution than that of a conventional AEC sensor. Thus, a part of theobject area can be assigned as reference of AEC. A pixel for use in AECis chosen from the pixels in the image capturing field 41, so AEC isappropriately carried out even if the size and shape of the body portionis changed. Moreover, since the minimum-value pixel having the lowestpixel value is used for AEC, the necessary dose is certainly applied tothe entire object area, and hence the image quality of the entire objectarea is improved.

Using the pixel provided in the image capturing field 41 for AEC caneliminate the need for providing the AEC sensor independent of the FPD26, and hence simplify the structure of the apparatus. In the case ofproviding the AEC sensor independent of the FPD 26 in front of the imagecapturing field 41 of the FPD 26, the AEC sensor attenuates the X-raysapplied to the FPD 26, but the present invention is free from such anattenuation problem.

In the above embodiment, the short pixels 55 and the normal pixels 40have approximately the same structure and the same sensitivity to theX-rays, and therefore the pixel values of the normal pixels 40 areestimated with high accuracy based on the pixel values of the shortpixels 55. This facilitates improving the accuracy of AEC. Also, thesame or similar structure of the pixels can ease manufacturing andreduce manufacturing costs.

In the above embodiment, the maximum-value pixel is chosen in additionto the minimum-value pixel. When the second integrated value of themaximum-value pixel has reached the second threshold value, the X-rayemission is stopped even if the first integrated value of theminimum-value pixel has not reached the first threshold value.Therefore, it is possible to prevent the excessive X-ray exposure in theentire object area.

In the above embodiment, before setting the minimum-value andmaximum-value pixels, the irradiation field is determined in the imagecapturing field 41, and the object area is determined in the irradiationfield by excluding the directly exposed area and the implant area. Theminimum-value and maximum-value pixels are chosen from the pixels of theobject area. This allows setting the minimum-value and maximum-valuepixels appropriately in the object area.

In the above embodiment, both the directly exposed area and the implantarea are excluded from the irradiation field to determine the objectarea, and the minimum-value and maximum-value pixels are set in theobject area. However, only the implant area may be excluded from theirradiation field, and the minimum-value and maximum-value pixels may beset in an area including the directly exposed area and the object area.In this case, the maximum-value pixel is probably chosen from the pixelsin the directly exposed area. Using this maximum-value pixel, a maximumX-ray dose applied to the irradiation field is checked, and therefore itis possible to prevent the excessive X-ray exposure of the patient.

Without determination of the object area, an index area may be assignedin advance in the image capturing field 41, and the minimum-value andmaximum-value pixels may be chosen from the pixels in the index area. Ina case that the position, size, shape, and the like of areas where tolocate the minimum-value and maximum-value pixels within the object areaare roughly known, the assignment of the index area can eliminate theneed for determining the irradiation field, the object area, and thelike. The exposure controller 32 can determine the minimum-value andmaximum-value pixels in the index area, so it is possible to easecalculation processing and accelerate processing speed for determiningthe minimum-value and maximum-value pixels.

In addition to the index area, an interest area may be assigned. In thiscase, for example, the exposure controller 32 sets the minimum-valuepixel in the index area, and the maximum-value pixel in the interestarea. This is effective when the X-ray transmittance is higher in theinterest area than in the other areas in the object area, as in the caseof the chest radiography described above. For example, in FIG. 14, thearea 73 a including the lung fields is assigned as the interest area,while the area 73 b including the mediastinum and the diaphragm havingthe lower X-ray transmittance than that of the lung fields is assignedas the index area. This brings about the same effect as the aboveembodiment. Furthermore, eliminating the need for determining the objectarea increases the processing speed.

The index area and the interest area are assigned by the operation ofthe console 14, for example. An area setting screen that schematicallyshows the image capturing field 41 is displayed on the monitor 75 of theconsole 14, and an arbitrary area is assigned on the screen as the indexarea and the interest area. Data of the assigned index area and interestarea is transmitted to the exposure controller 32 of the electroniccassette 13.

In the above embodiment, both of the minimum-value and maximum-valuepixels are determined and AEC is performed based on the first and secondintegrated values. However, only the minimum-value pixel is determinedand AEC may be performed based on only the first integrated value. IfAEC is performed based on only the first integrated value, the necessarydose is certainly applied, so AEC is carried out appropriately. As forprevention of the excessive X-ray exposure, instead of determination ofthe maximum-value pixel, for example, maximum emission time may bedetermined and the X-ray emission may be forcefully stopped when a lapseof the maximum emission time has counted by a timer. As a matter ofcourse, in a method of setting the maximum-value pixel, the X-ray doseapplied to the object area is actually measured. Thus, the method ofsetting the maximum-value pixel prevents the excessive X-ray exposuremore effectively than the method of setting the maximum emission time.

In the above embodiment, the short pixel 55 that is directly connectedto the signal line 49 is used as the detection pixel for detecting theX-rays. However, another type of detection pixel that is connected tothe signal line 49 through the TFT being the switching element, as withthe normal pixel 40, may be provided instead. Using this type ofdetection pixel allows control of charge accumulation time of thedetection pixel, and readout of the pixel value Vout at arbitrarytiming.

In the above embodiment, the short pixels 55 are arranged together withthe normal pixels 40, and the pixel values of the normal pixels 40 areestimated based on the pixel values Vout of the short pixels 55.However, instead of the short pixels 55, X-ray sensors that functionsimilarly to the short pixels 55 may be provided such that each X-raysensor is disposed between the normal pixels 40 adjoining to each other,in order to estimate the pixel values of the normal pixels 40 based ondose detection values of the X-ray sensors. The X-ray image read outafter completion of the X-ray emission has a defect caused by the shortpixels 55, which are dealt with as defective pixels. However,disposition of the X-ray sensors between the normal pixels 40 caneliminate the need for providing the short pixels 55 being the defectpixels, and therefore ease the defect correction.

For the purpose of improving accuracy in the defect correction forcorrecting the effect of the short pixels 55 being the detection pixels,an FPD 90 shown in FIG. 17 is composed of two types of combined pixels91 and 92, instead of the normal pixels 40 and the short pixels 55. Thecombined pixel 91 includes two subpixels 93 and 94. The combined pixel92 includes two subpixels 93 and 96.

The subpixels 93 and 94 are specifically used for image detection aswith the normal pixel 40, while the subpixel 96 is used for AEC as withthe detection pixel such as the short pixel 55. Thus, the combined pixel91 is constituted of the two subpixels 93 and 94 for use in the imagedetection, while the combined pixel 92 is constituted of the subpixel 93for use in the image detection and the subpixel 96 functioning as thedetection pixel. Each of the combined pixels 91 and 92 is approximatelythe same size as the single normal pixel 40. Each of the subpixels 93,94, and 96 is approximately half of the single normal pixel 40 in size.The combined pixels 92 are distributed over the entire image capturingfield 41 at an appropriate rate, as with the short pixels 55.

Each of the subpixels 93, 94, and 96 is made of a photodiode. In thecombined pixel 91, the subpixels 93 and 94 are connected to the signalline 49 through the TFT 46 in parallel. In the combined pixel 92, on theother hand, the subpixel 93 is connected to the signal line 49 throughthe TFT 46, and the subpixel 96 is directly connected to the signal line49 without through the TFT 46 just as with the short pixel 55.

In reading out the X-ray image, a sum of electric charge accumulated inthe two subpixels 93 and 94 is read out from the combined pixel 91. Fromthe combined pixel 92, only electric charge accumulated in the subpixel93 is read out. The amount of electric charge accumulated in a subpixelis proportional to the size of the subpixel. Therefore, if the combinedpixels 91 and 92 are applied with the same X-ray dose, the amount of theelectric charge read out from the combined pixel 92 is approximatelyhalf of that from the combined pixel 91. Since the subpixel 96 isdirectly connected to the signal line 49, the electric charge producedin the subpixel 96 continuously flows into the signal line 49. Theelectric charge that flows from the subpixel 96 is detected as the pixelvalue Vout for use in AEC.

In the defect correction of the X-ray image, for example, a pixel valueof the combined pixel 92 is doubled. In other words, the pixel value ofthe combined pixel 92 is multiplied by a coefficient that is calculatedin advance based on the ratio in size between the subpixel 93 of thecombined pixel 92 and the two subpixels 93 and 94 of the combined pixel91. The defect correction is necessary even with the use of the combinedpixels 92, owing to provision of the subpixels 96, which do notcontribute detection of the X-ray image. However, as compared with thecase of providing the short pixels 55, correction accuracy is improveddue to the subpixels 93. For this reason, deterioration in the X-rayimage is prevented, when compared with the above embodiment using theshort pixels 55.

In the above embodiment, the pixel values of the normal pixels 40 areestimated based on the pixel values Vout of the short pixels 55 beingthe detection pixels, and the minimum-value and maximum-value pixelsused in AEC are chosen from the normal pixels 40 based on the estimatedpixel values. However, the minimum-value and maximum-value pixels usedin AEC may be chosen from the detection pixels. In this case, the firstand second integrated values are obtained based on the pixel values Voutof the detection pixels. Since the detection pixels are distributed overthe entire image capturing field 41 at the proper rate, AEC isappropriately carried out irrespective of the body portion to be imaged,even if the detection pixels and their pixel values Vout are directlyused in AEC. The number of the detection pixels is less than that of thenormal pixels 40, as a matter of course. Therefore, spatial positionalaccuracy deteriorates as compared with the case of choosing the pixelsfor use in AEC from the normal pixels 40, but accuracy in the pixelvalues themselves is improved because the measured pixel values are usedinstead of the estimated pixel values.

The FPD has the photodiodes and the TFTs formed in a glass substrate inthis embodiment, but a CMOS (complementary metal oxide semiconductor)type FPD may be used instead. The CMOS type FPD has an image capturingfield having an arrangement of pixels each of which is constituted of aphotodiode and a switching element formed in a silicon substrate. TheCMOS type FPD can perform so-called non-destructive readout in whichwhile a pixel keeps holding electric charge, a voltage corresponding tothe accumulated electric charge is read out from the pixel. Thus, everypixel is available as both the normal pixel for use in image detectionand the detection pixel for use in AEC. In the case of using the CMOStype FPD, the exposure controller 32 performs determination of theobject area, choice of the minimum-value and maximum-value pixels, andcalculation of the first and second integrated values based on themeasured pixel values, without estimating the pixel values.

In the above embodiment, the minimum-value pixel, which outputs thelowest pixel value, is set as the typical low-value pixel for use in AECto obtain the X-ray image with favorable image quality. However, theminimum-value pixel may not be necessarily assigned as the typicallow-value pixel, as long as the typical low-value pixel is determinedfrom a plurality of low-value pixels including the minimum-value pixel.The plurality of low-value pixels refer to pixels that output apredetermined range of pixel values including the lowest pixel value,and more specifically, pixels that output pixel values within a range oflowest 10% to 20% in an entire range from the lowest pixel value to thehighest pixel value. Taking the histogram of FIG. 12 as an example, outof the normal pixels 40 located within the object area, a plurality ofnormal pixels 40 whose estimated pixel values are within the range oflowest 10% to 20% are denoted as the low-value pixels.

The pixel determiner 58 determines the typical low-value pixel based onthe histogram from the plurality of low-value pixels. As an example ofdetermining the typical low-value pixel from the low-value pixelsexcluding the minimum-value pixel, for example, a low-value pixel thatoutputs a mean or median value may be determined as the typicallow-value pixel. According to this method, even if the minimum-valuepixel outputs an abnormal pixel value, the minimum-value pixel is notdetermined as the typical low-value pixel, so AEC is carried outappropriately. The pixel value of the determined typical low-value pixelis integrated, as in the case of the above embodiment, and an integratedvalue is used as the first integrated value.

A plurality of typical low-value pixels may be determined from theplurality of low-value pixels. In this case, for example, the pixeldeterminer 58 chooses one of the low-value pixels in the image capturingfield 41, and then determines as the typical low-value pixels a pixelgroup that is composed of the chosen pixel and a plurality of low-valuepixels around the chosen pixel. In a case where there are a plurality oftypical low-value pixels, for example, a mean, median, or sum of thepixel values outputted from the typical low-value pixels is determined.The mean, median, or sum value is integrated and used as the firstintegrated value. The first threshold value to be compared with thefirst integrated value is appropriately determined in accordance withthe type (mean, median, or sum) of the first integrated value.

The same goes for the typical high-value pixel used for prevention ofthe excessive X-ray exposure. To be more specific, the maximum-valuepixel, which outputs the highest pixel value, is set as the typicalhigh-value pixel in the above embodiment. However, the maximum-valuepixel may not be necessarily assigned as the typical high-value pixel,as long as the typical high-value pixel is determined from a pluralityof high-value pixels including the maximum-value pixel. The plurality ofhigh-value pixels refer to pixels that output a predetermined range ofpixel values including the highest pixel value, and more specifically,pixels that output pixel values within a range of highest 10% to 20% inthe entire range from the lowest pixel value to the highest pixel value.Taking the histogram of FIG. 12 as an example, out of the normal pixels40 located within the object area, a plurality of normal pixels 40 whoseestimated pixel values are within the range of highest 10% to 20% aredenoted as the high-value pixels.

The pixel determiner 58 determines the typical high-value pixel based onthe histogram from the plurality of high-value pixels. The pixel valueof the determined typical high-value pixel is integrated, as in the caseof the above embodiment, and an integrated value is used as the secondintegrated value. According to this method, even if the maximum-valuepixel outputs an abnormal pixel value, the maximum-value pixel is notdetermined as the typical high-value pixel, so AEC is carried outappropriately. A plurality of typical high-value pixels may bedetermined, as in the case of the typical low-value pixels. In thiscase, for example, a mean, median, or sum of the pixel values outputtedfrom the plurality of typical high-value pixels is determined. The mean,median, or sum value is integrated and used as the second integratedvalue. The second threshold value to be compared with the secondintegrated value is appropriately determined in accordance with the type(mean, median, or sum) of the second integrated value.

Note that, how to determine the typical low-value and high-value pixelsis not limited to above, and any method is usable as long as arelatively low-value pixel and a relatively high-value pixel in theimage capturing field 41 are determined as the typical low-value andhigh-value pixels. For example, in order to prevent a defective pixel,which outputs an abnormal pixel value, from being set as the typicallow-value or high-value pixel, the minimum-value pixel and themaximum-value pixel may be excluded from the low-value pixels and thehigh-value pixels, respectively. The typical low-value pixel isdetermined from the low-value pixels excluding the minimum-value pixel,and the typical high-value pixel is determined from the high-valuepixels excluding the maximum-value pixel. In the case of determiningboth the typical high-value and low-value pixels, the most importantmatter is that AEC is performed with referring to both the relativelylow-value and high-value pixels in the object area, for the purpose ofobtaining the favorable image quality and preventing the excessive X-rayexposure. A slight difference in how to concretely determine the typicallow-value and high-value pixels is inessential and has no influence onan effect of the present invention, though it only causes a smalldifference in the pixel values from the determined typical low-value andhigh-value pixels.

The electronic cassette 13 and the console 14 are wirelessly connectedin the above embodiment, but may be connected through a wire. Theconsole 14 and the electronic cassette 13 are separate in the aboveembodiment, but the console 14 may not be necessarily independent. Theelectronic cassette 13 may have the function of the console 14.Alternatively, another imaging control device specific to the control ofthe electronic cassette 13 may be provided between the electroniccassette 13 and the console 14, and the console 14 may take charge ofonly easy functions e.g. input of the imaging condition and display ofthe X-ray image. The console 14 and the source controller 17 may beintegrated into one unit. The present invention may be applied to aninstalled type of X-ray image detecting device in which the FPD iscontained in the imaging stand, instead of the electronic cassette beinga portable type of X-ray image detecting device.

The present invention is applicable to a radiation imaging system usinganother type of radiation such as γ-rays instead of the X-rays.

Although the present invention has been fully described by the way ofthe preferred embodiment thereof with reference to the accompanyingdrawings, various changes and modifications will be apparent to thosehaving skill in this field. Therefore, unless otherwise these changesand modifications depart from the scope of the present invention, theyshould be construed as included therein.

What is claimed is:
 1. A radiation imaging apparatus comprising: animage detector for detecting a radiographic image of an object, saidimage detector including a plurality of pixels arranged in an imagecapturing field, each of said pixels receiving radiation emitted from aradiation source and outputting a pixel value in accordance with areceived radiation dose; a pixel determiner for determining at least onetypical low-value pixel from said pixels based on said pixel values, andsetting said typical low-value pixel as an exposure control pixel, saidpixel determiner determining said typical low-value pixel out of saidpixels present within an index area being predetermined in said imagecapturing field in accordance with a body portion to be imaged; and acomparator for comparing a first integrated value being an integratedvalue of said pixel value of said typical low-value pixel with apredetermined first threshold value, and performing radiation emissioncontrol such that, when said first integrated value has reached saidfirst threshold value, said radiation source stops emitting saidradiation.
 2. The radiation imaging apparatus according to claim 1,further comprising: an irradiation field determiner for determining anirradiation field based on said pixel values, said irradiation fieldbeing a field irradiated with said radiation in said image capturingfield, wherein said pixel determiner determines in said irradiationfield a directly exposed area being an area applied with said radiationdirectly without through said object, an implant area being an area ofan implant implanted in said object, and an object area being an areaexcluding said directly exposed area and said implant area from saidirradiation field; and said typical low-value pixel is determined out ofsaid pixels in said object area.
 3. The radiation imaging apparatusaccording to claim 2, wherein said pixel determiner determines saidobject area based on a histogram of said pixel values of said pixels insaid irradiation field.
 4. The radiation imaging apparatus according toclaim 3, wherein radiation absorptance is higher in said index area thanin said interest area.
 5. The radiation imaging apparatus according toclaim 1, wherein said pixels include a plurality of normal pixels forspecific use in detection of said radiographic image, and a plurality ofdetection pixels distributed throughout said image capturing field todetect said radiation dose.
 6. The radiation imaging apparatus accordingto claim 5, further comprising: a pixel value estimator for estimatingsaid pixel value of said normal pixel based on said pixel values of saiddetection pixels near said normal pixel to be estimated; and said pixeldeterminer determines said typical low-value pixel based on saidestimated pixel values.
 7. The radiation imaging apparatus according toclaim 6, wherein said image detector has a plurality of pixel groupseach including one or more of said normal pixels and one or more of saiddetection pixels, and said detection pixels are laid out differentlybetween said pixel groups adjoining to each other; and said pixel valueestimator estimates said pixel value of said normal pixel of a firstpixel group, based on said pixel value of said detection pixel belongingto said first pixel group and said pixel value of said detection pixelbelonging to a second pixel group adjoining to said first pixel group.8. The radiation imaging apparatus according to claim 5, wherein saidpixel determiner determines said typical low-value pixel out of saiddetection pixels.
 9. The radiation imaging apparatus according to claim5, wherein signal lines electrically connected to said pixels are routedin said image capturing field to output said pixel values; and saiddetection pixel is connected to said signal line directly or through aswitching element.
 10. The radiation imaging apparatus according toclaim 5, wherein said pixels include a combined pixel that is composedof a first subpixel functioning as said normal pixel and a secondsubpixel functioning as said detection pixel.
 11. A radiation imagingapparatus comprising: an image detector for detecting a radiographicimage of an object, the image detector including a plurality of pixelsarranged in an image capturing field, each of the pixels receivingradiation emitted from a radiation source and outputting a pixel valuein accordance with a received radiation dose; a pixel determiner fordetermining at least one typical low-value pixel from the pixels basedon the pixel values, and setting the typical low-value pixel as anexposure control pixel; and a comparator for comparing a firstintegrated value being an integrated value of the pixel value of thetypical low-value pixel with a predetermined first threshold value, andperforming radiation emission control such that, when the firstintegrated value has reached the first threshold value, the radiationsource stops emitting the radiation, wherein the pixel determinerdetermines a minimum-value pixel that outputs a lowest pixel value basedon a histogram of the pixel values, and sets the minimum-value pixel asthe typical low-value pixel.
 12. The radiation imaging apparatusaccording to claim 11, further comprising: an irradiation fielddeterminer for determining an irradiation field based on said pixelvalues, said irradiation field being a field irradiated with saidradiation in said image capturing field, wherein said pixel determinerdetermines in said irradiation field a directly exposed area being anarea applied with said radiation directly without through said object,an implant area being an area of an implant implanted in said object,and an object area being an area excluding said directly exposed areaand said implant area from said irradiation field; and said typicallow-value pixel is determined out of said pixels in said object area.13. The radiation imaging apparatus according to claim 12, wherein saidpixel determiner determines said object area based on a histogram ofsaid pixel values of said pixels in said irradiation field.
 14. Theradiation imaging apparatus according to claim 13, wherein radiationabsorptance is higher in said index area than in said interest area. 15.The radiation imaging apparatus according to claim 11, wherein saidpixels include a plurality of normal pixels for specific use indetection of said radiographic image, and a plurality of detectionpixels distributed throughout said image capturing field to detect saidradiation dose.
 16. The radiation imaging apparatus according to claim15, further comprising: a pixel value estimator for estimating saidpixel value of said normal pixel based on said pixel values of saiddetection pixels near said normal pixel to be estimated; and said pixeldeterminer determines said typical low-value pixel based on saidestimated pixel values.
 17. The radiation imaging apparatus according toclaim 16, wherein said image detector has a plurality of pixel groupseach including one or more of said normal pixels and one or more of saiddetection pixels, and said detection pixels are laid out differentlybetween said pixel groups adjoining to each other; and said pixel valueestimator estimates said pixel value of said normal pixel of a firstpixel group, based on said pixel value of said detection pixel belongingto said first pixel group and said pixel value of said detection pixelbelonging to a second pixel group adjoining to said first pixel group.18. The radiation imaging apparatus according to claim 15, wherein saidpixel determiner determines said typical low-value pixel out of saiddetection pixels.
 19. The radiation imaging apparatus according to claim15, wherein signal lines electrically connected to said pixels arerouted in said image capturing field to output said pixel values; andsaid detection pixel is connected to said signal line directly orthrough a switching element.
 20. The radiation imaging apparatusaccording to claim 15, wherein said pixels include a combined pixel thatis composed of a first subpixel functioning as said normal pixel and asecond subpixel functioning as said detection pixel.